The present invention relates generally to the field of optical coherence tomography and, more particularly, to systems and methods for multi-channel optical coherence tomography.
Modern instrumental analysis techniques include sophisticated optical scanning at ever-greater resolutions. For example, the interferometric technique, optical coherence tomography (“OCT”) provides three-dimensional images of scattering samples in the micrometer resolution range. Modern OCT techniques include swept-source OCT (“SS-OCT”) and spectral-domain OCT (“SD-OCT”), which are commonly used in biomedical imaging systems. An exemplary SS-OCT approach is described by S. R. Chinn, et al. in “Optical coherence tomography using a frequency-tunable optical source,” Opt. Lett., Vol. 22, Issue 5 (1997), pp. 340-342. An exemplary SD-OCT approach is described by A. F. Fercher, et al. in “Measurement of intraocular distances by backscattering spectral interferometry,” Opt. Comm's, Vol. 117, Issues 1-2 (1995), pp. 43-48. For ease and brevity of explanation, the references cited herein are provided as citations. However, all references cited herein should be considered incorporated by reference in their entirety.
Broadly, both SS-OCT and SD-OCT generate a spectral interference signal between a reference signal and a sample signal using a low-coherence light source. In SS-OCT, the light source is swept in wavelength. In SD-OCT, the light source has a broadband spectrum. The spectral interference signal is a function of the wavelength of the light source illuminating the sample. The spectral interference signal Fourier transform amplitude gives the sample refractive index distribution along the sample depth, customarily referred to as the z-direction. Both SS-OCT and SD-OCT obtain volumetric data by scanning the sample surface along the fast lateral direction (customarily referred to as the x-direction) and the slow lateral direction (customarily referred to as the y-direction).
An additional technique, Doppler OCT, employs either SS-OCT or SD-OCT to determine the velocity of moving particles inside the sample, in addition to the structural image, by computing the phase difference of the OCT signal between two instants at a given location. One of ordinary skill in the art will understand that the maximum absolute value of the phase is π and its minimum absolute value is limited by the phase noise, a result noted by B. H. Park, et al., in “Real-time fiber-based multi-functional spectral-domain optical coherence tomography at 1.3 μm,” Opt. Express, Vol. 13, No. 11 (2005), pp. 3931-3944. One of ordinary skill in the art will also understand that the measurable velocity is proportional to the phase difference and inversely proportional to the time interval between the two measurements. Conventional Doppler OCT techniques, such as that presented in U.S. Pat. No. 6,549,801, execute phase difference measurements between adjacent lateral locations of the fast scanning direction (x-direction). An additional example is provided by Y. Zhao, et al., in “Doppler standard deviation imaging for clinical monitoring of in vivo human skin blood flow,” Opt. Lett., Vol. 25, No. 18 (2000), pp. 1358-1360.
One of ordinary skill in the art will understand that measurement of adjacent lateral locations requires dense scanning in order to scan locations close enough together to return correlated OCT signals. This correlation requirement complicates the process of obtaining densely collected measurements. For example, ophthalmological and diagnostic applications commonly employ Doppler OCT to obtain motion information for retinal or choroidal blood vessels. However, typical modern in vivo ophthalmologic OCT measurements require a short volume acquisition time due to motion artifacts, power exposure limits, and other safety and environmental considerations.
Even with these limitations, recent developments in SS-OCT and SD-OCT have achieved scanning speeds over 400 kHz. Exemplary techniques are shown by T. Klein, et al., in “Megahertz OCT for ultrawide-field retinal imaging with a 1050 nm Fourier domain mode-locked laser,” Opt. Express, Vol. 19, No. 4 (2011), pp. 3044-3062, by B. Potsaid et al., in “Ultrahigh speed 1050 nm swept source/Fourier domain OCT retinal and anterior segment imaging at 100,000 to 400,000 axial scans per second,” Opt. Express, Vol. 18, No. 19 (2010), pp. 20029-20048, and by L. An, et al., in “Optical microangiography provides correlation between microstructure and microvasculature of optic nerve head in human subjects,” J. Biomed. Opt., Vol. 17, No. 11 (2012), pp. 116018-1-116018-6.
A person of ordinary skill in the art will understand that SS-OCT Doppler measurements require accurate synchronization between the detected signal and the sweeping scan illuminations. One approach to this synchronization problem is to use a fiber Bragg grating, as described by H. C. Hendargo, et al. in “Doppler velocity detection limitations in spectrometer-based versus swept-source optical coherence tomography,” Biomed. Opt. Express, Vol. 2, No. 8 (2011), pp. 2175-2188.
SS-OCT systems configured accordingly can be employed to obtain high-resolution volume information with a short time interval between adjacent sample point locations. So configured, standard Doppler OCT techniques can be employed to measure relatively fast motion occurring inside the sample. This approach has been used to measure blood flow near the optic nerve head in a human retina by W. Choi, et al., as described in “Phase-sensitive swept-source optical coherence tomography imaging of the human retina with a vertical cavity surface-emitting laser light source,” Opt. Lett., Vol. 38, Issue 3 (2013), pp. 338-340.
However, this approach is less effective in measuring relatively slow motion occurring inside the sample. One of ordinary skill in the art will understand that sensitivity to slower motion within the sample ordinarily requires decreasing the time interval between phase measurements. One alternative approach is to modify the scanning protocol in the fast scanning direction (x-direction). Another alternative approach is to compute the phase difference between measurements along the slow scanning direction (y-direction), as proposed by B. Braaf, et al., in “Angiography of the retina and the choroid with phase-resolved OCT using interval-optimized backstitched B-scans,” Opt. Express, Vol. 20, No. 18 (2012), pp. 20516-20534. These alternative approaches, however, fail to overcome the trade-off between spatial resolution (or scanning range size) and measurable velocity range. As such, several additional approaches have been proposed.
One such approach uses two spatially-separated probing beams. An exemplary expression of this approach is shown in U.S. Pat. Pub. No. 2012/0120408 A1. Generally, this approach scans a sample with two probing beams at the same speed, with a constant spatial shift between the two probing beams. In typical systems, this approach achieves improved motion sensitivity without excessive reduction in spatial resolution or lateral scanning range.
In a common dual-beam example, each probing beam scans the same location at different times. One probing beam (the “following” beam) scans the target location after a constant delay T with respect to the other probing beam (the “preceding” beam). For each target location, the phase of the preceding beam signal acquired at time t0 is compared with the phase of the following beam signal acquired at time t0+T. The common dual-beam approach allows for some measurement of relatively slower particle motions.
This approach has been successfully employed to obtain angiographies of the human choroid, as noted by F. Jaillon, et al., in “Variable velocity range imaging of the choroid with dual-beam optical coherence angiography,” Opt. Express, Vol. 20, No. 1 (2012), pp. 385-396. Additionally, the measurable velocity range can be adjusted by modifying the delay, T, between the two probing beams, to optimize vessel contrast. As the delay, T, between the beams increases, slower motion can be measured. Different implementations of this technique have been demonstrated.
The original implementation of the common dual-beam technique used polarization multiplexing to generate the distinct probing beams, which introduced operational challenges and complexity. For example, the polarization multiplexing implementation requires the interferometer to use polarization-maintaining fibers and the probing arm to use a Wollaston prism or polarization beam splitter in order to separate the two beams.
As such, this implementation may be sensitive to the intrinsic birefringence of the probed sample, which reduces vessel contrast. This sensitivity to birefringence can be reduced somewhat by adding a Faraday rotator and a quarter waveplate in the probing arm, as described by M. Makita, et al., in “Dual-beam-scan Doppler optical coherence angiography for birefringence-artifact-free vasculature imaging,” Opt. Express, Vol. 20, No. 3 (2012), pp. 2681-2692. However, this approach increases system complexity and cost.
Another implementation uses a second light source and combines the two probing beams with a non-polarizing beam splitter, as described by S. Zotter, et al., in “Visualization of microvasculature by dual-beam phase-resolved Doppler optical coherence tomography,” Opt. Express, Vol 19, No. 2 (2011), pp. 1217-1227. This second implementation generates the two probing beams from a beam splitter and two independent light sources. One shortcoming of this implementation is that the beam splitter causes significant signal losses of the retinal backscattered signal. Additionally, this approach also increases system complexity and cost.
Another approach in some dual-beam systems is to employ two angularly separated probing beams. For example, such systems use two probing beams impinging on the same sample location at different angles. This approach is shown by, for example, R. Werkmeister, et al., in “Bidirectional Doppler Fourier-domain optical coherence tomography for measurement of absolute flow velocities in human retinal vessels,” Opt. Lett., Vol. 33, Issue 24 (2008), pp. 2967-2969. Generally, this approach supports measuring the absolute velocity of motion inside the sample.
An early demonstration of application of this approach to laser Doppler velocimetry is presented in U.S. Pat. No. 5,900,928. Conventional Doppler OCT only determines the projection of the velocity direction over the probing beam direction or the axial velocity. Further, typical bidirectional dual-beam approaches can only determine absolute velocity under certain conditions. In particular, the typical approach requires that the plane defined by the two probing beams also contains the direction of the velocity vector under measurement. One approach to meeting this requirement is the rotation of the bulk optics to achieve alignment between the dual-beam plane and the desired velocity direction, as demonstrated by C. Blatter, et al., in “Dove prism based rotating dual beam bidirectional Doppler OCT,” Biomed. Opt. Express, Vol. 4, No. 7 (2013), pp. 1188-1203. In this technique, alignment is achieved by rotating bulk optics, in particular a Dove prism, to measure blood velocity in the optic nerve head of the eye. One of ordinary skill in the art will appreciate that this technique tends to increase the cost and complexity of systems using this approach.
Therefore, there is a need for a system and/or method for optical coherence tomography that addresses at least some of the problems and disadvantages associated with conventional systems and methods.